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Abstract:

Methods and apparatus for operating an MRI system is provided. The
disclosure provides a diffusion-prepared driven-equilibrium preparation
for an imaging volume and acquiring 3-dimensional k-space data from said
prepared volume by a plurality of echoplanar readouts of stimulated
echoes. An excitation radio-frequency signal and first and second
inversion RF signals are provided to define a field-of-view (FOV).

Claims:

1. A method for interleaved MR imaging, comprising: providing an
excitation radio-frequency (RF) signal; providing first and second
inversion RF signals to define a field-of-view (FOV).

2. The method of claim 1, wherein said first and second inversion RF
signals comprise first and second inversion RF pulses.

3. The method of claim 2, wherein said first inversion RF pulse is
applied substantially immediately after said excitation RF signal.

4. The method of claim 2, wherein said first and second inversion RF
pulses are separated by a time approximately 5 ms or larger.

5. The method of claim 2, wherein said first and second inversion RF
pulses are separated by a time less than about 5 ms.

6. The method of claim 2, wherein said first and second inversion RF
pulses are separated by a time more than about 5 ms.

7. The method of claim 2, further comprising providing slice-selective
gradients that are selected such that magnetization within said FOV is
substantially preserved while magnetization external to said FOV is
substantially suppressed, thereby allowing magnetization in each of a
plurality of slices to be substantially maintained in its equilibrium
state while exciting and imaging one or more others of said plurality of
slices.

8. A control system for an MRI apparatus, comprising: a control component
configured to generate one or more instructions for: providing an
excitation radio-frequency (RF) signal; providing first and second
inversion RF signals to define a field-of-view (FOV).

9. The system of claim 8, wherein said first and second inversion RF
signals comprise first and second inversion RF pulses.

10. The system of claim 9, wherein said first inversion RF pulse is
applied substantially immediately after said excitation RF signal.

11. The system of claim 9, wherein said first and second inversion RF
pulses are separated by a time approximately 5 ms or larger.

12. The system of claim 9, wherein said first and second inversion RF
pulses are separated by a time less than about 5 ms.

13. The system of claim 9, wherein said first and second inversion RF
pulses are separated by a time more than about 5 ms.

14. The system of claim 9, wherein said one or more instructions further
includes an instruction for providing slice-selective gradients that are
selected such that magnetization within said FOV is substantially
preserved while magnetization external to said FOV is substantially
suppressed, thereby allowing magnetization in each of a plurality of
slices to be substantially maintained in its equilibrium state while
exciting and imaging one or more others of said plurality of slices.

15. A method for correcting a motion artifact during MR imaging,
comprising: acquiring navigation data substantially together with imaging
data; determining whether to re-acquire said imaging data based on said
navigation data; and re-acquiring said imaging data based on said
determination.

16. The method of claim 15, wherein said motion artifact is due to
intra-shot motion.

17. The method of claim 16, wherein said motion artifact is due to
inter-shot motion.

18. The method of claim 15, wherein said determining and re-acquiring are
performed substantially real-time.

19. The method of claim 15, wherein said navigation data comprises 2D
k-space navigation echoes, and wherein said determining comprises
identifying value and position of the largest signal in said 2D k-space
to see if either of said value or position is outside of a corresponding
selected range.

20. The method of claim 15, wherein said MR imaging comprises a
multi-average singleshot EPI operated as at least one of DWI, DTI, and
fMRI.

21. The method of claim 15, wherein said MR imaging comprises at least
one of spin-echo, multiple spin-echo, gradient-echo, and segmented
gradient-echo.

22. A system for correcting a motion artifact during MR imaging,
comprising: a control component configured to generate one or more
instructions for: acquiring navigation data substantially together with
imaging data; determining whether to re-acquire said imaging data based
on said navigation data; and re-acquiring said imaging data based on said
determination.

23. The system of claim 22, wherein said motion artifact is due to
intra-shot motion.

24. The system of claim 23, wherein said motion artifact is due to
inter-shot motion.

25. The system of claim 22, wherein said determining and re-acquiring are
performed substantially real-time.

26. The system of claim 22, wherein said navigation data comprises 2D
k-space navigation echoes, and wherein said determining comprises
identifying value and position of the largest signal in said 2D k-space
to see if either of said value or position is outside of a corresponding
selected range.

27. The system of claim 22, wherein said MR imaging comprises a
multi-average singleshot EPI operated as at least one of DWI, DTI, and
fMRI.

28. The system of claim 22, wherein said MR imaging comprises at least
one of spin-echo, multiple spin-echo, gradient-echo, and segmented
gradient-echo.

Description:

CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This is a divisional patent application of U.S. patent application
Ser. No. 11/732,382 filed Apr. 2, 2007, which claims the benefit under 35
U.S.C. §119(e) of U.S. Provisional Patent Application No.
60/788,533, filed on Mar. 31, 2006, and titled "SYSTEMS AND METHODS FOR
MAGNETIC RESONANCE IMAGING," the entirety of which is hereby incorporated
by reference.

BACKGROUND OF THE INVENTION

[0003] 1. Field of the Invention

[0004] This application relates to magnetic resonance imaging.

[0005] 2. Description of the Related Art

[0006] Diffusion-weighted magnetic resonance (MR) imaging (DWI) is a known
tool for detecting abnormal water diffusion in the brain (e.g., ischemic
stroke). The directional information obtained using diffusion tensor MRI
(DTI) is valuable in understanding as well as evaluating white matter
abnormalities in neurological diseases, such as Alzheimer disease,
schizophrenia, multiple sclerosis, and neurofibromatosis. DWI and DTI may
also give useful information about the development and disorders of
ordered structures in extracranial organs such as the heart, kidney,
breast, and prostate.

[0007] Although DTI can provide useful information about white-matter
diseases in the brain, high resolution DTI of brain regions near the
temporal bone or sinuses, of small neural structures such as the spinal
cord or optic nerve, or of extracranial organs in vivo has been difficult
to achieve using conventional two-dimensional (2D) singleshot
diffusion-weighted EPI techniques (2D ss-DWEPI). There are strong
non-uniform local magnetic fields created by the magnetic susceptibility
changes at tissue/bone or tissue/air interfaces, which typically induce
severe distortion on the resultant ss-DWEPI images. The amount of
susceptibility induced geometric distortion is proportional to the total
sampling time in EPI. Typically, increasing spatial resolution requires
an increase in the duration of the data acquisition window, which in turn
increases the distortion from off-resonance effects. As a result, the
spatial resolution obtained using conventional 2D ss-EPI is generally
much lower than that obtainable with conventional, multi-shot MRI, giving
decreased resolution for measurements of interest, such as white matter
tract anatomy and nerve fiber anatomy. For these reasons, 2D ss-DWEPI has
been clinically useful only for moderately low resolution intracranial
applications. EPI with parallel imaging has been successfully applied to
high-resolution brain DWI and DTI studies resulting in substantial image
quality improvement.

[0008] There are several non-EPI singleshot DWI techniques, which include
multiple spin-echo sequences (e.g., ss-FSE (or HASTE) and GRASE), STEAM,
and fast gradient echo sequences (FLASH), that complete the total data
acquisition following a single diffusion weighting. These 2D sequences
typically acquire slightly more than half of the ky encodings in
about 500 ms after a single diffusion weighting preparation. These
non-EPI singleshot techniques typically employ relatively thick slices to
overcome their intrinsic low SNR.

[0009] Multishot imaging techniques may be used to increase SNR, improve
spatial resolution and reduce susceptibility induced artifacts. However
most multishot DWI acquisition techniques suffer from the instability of
phase errors among shots due to global or localized motions during
application of the large diffusion gradients. There has been reasonable
success with techniques that use navigator echoes to detect and correct
phase errors, or that use non-singleshot-EPI approaches that are less
sensitive to phase errors. Because most of these are 2D acquisition
techniques, they produce relatively poor resolution along the slice
direction.

SUMMARY OF THE INVENTION

[0010] Certain embodiments of the present disclosure relate to methods and
apparatus for operating an MRI system. The disclosure provides a
diffusion-prepared driven-equilibrium preparation for an imaging volume,
and acquiring 3-dimensional k-space data from the prepared volume by a
plurality of echoplanar readouts of stimulated echoes.

[0011] In certain embodiments, the diffusion-prepared driven-equilibrium
preparation includes a single diffusion-prepared driven-equilibrium
preparation. In certain embodiments, the 3-dimensional k-space data
includes a raw data that has not been transformed. In certain
embodiments, the MRI system is operated for diffusion-weighted MR imaging
(DWI). In certain embodiments, the MRI system is operated for diffusion
tensor MR imaging (DTI). In certain embodiments, the acquiring of
3-dimensional k-space data includes acquiring substantially entire
3-dimensional k-space data.

[0012] Certain embodiments of the present disclosure relate to a method
for interleaved MR imaging. The method includes providing an excitation
radio-frequency (RF) signal, and providing first and second inversion RF
signals to define a field-of-view (FOV).

[0013] In certain embodiments, the first and second inversion RF signals
include first and second inversion RF pulses. In certain embodiments, the
first inversion RF pulse is applied substantially immediately after the
excitation RF signal. In certain embodiments, the first and second
inversion RF pulses are separated by a time approximately 5 ms or larger.
In certain embodiments, the method further includes providing
slice-selective gradients that are selected such that magnetization
within the FOV is substantially preserved while magnetization external to
the FOV is substantially suppressed, thereby allowing magnetization in
each of a plurality of slices to be substantially maintained in its
equilibrium state while exciting and imaging one or more others of the
plurality of slices.

[0014] Certain embodiments of the present disclosure relate to a method
for correcting a motion artifact during MR imaging. The method includes
acquiring navigation data substantially together with imaging data. The
method further includes determining whether to re-acquire the imaging
data based on the navigation data. The method further includes
re-acquiring the imaging data based on the determination.

[0015] In certain embodiments, the motion artifact is due to intra-shot
motion. In certain embodiments, the motion artifact is due to inter-shot
motion. In certain embodiments, the determining and re-acquiring are
performed substantially real-time. In certain embodiments, the navigation
data includes 2D k-space navigation echoes, and the determining includes
identifying value and position of the largest signal in the 2D k-space to
see if either of the value or position is outside of a corresponding
selected range. In certain embodiments, the MR imaging includes a
multi-average singleshot EPI operated as at least one of DWI, DTI, and
fMRI. In certain embodiments, the MR imaging includes at least one of
spin-echo, multiple spin-echo, gradient-echo, and segmented
gradient-echo.

[0016] Certain embodiments of the present disclosure relate to a control
system for an MRI apparatus. The system includes a control component
configured to generate one or more instructions for providing
diffusion-prepared driven-equilibrium preparation for an imaging volume,
and acquiring a substantially entire 3-dimensional k-space data from the
prepared volume by a plurality of echoplanar readouts of stimulated
echoes.

[0017] In certain embodiments, the diffusion-prepared driven-equilibrium
preparation includes a single diffusion-prepared driven-equilibrium
preparation. In certain embodiments, the 3-dimensional k-space data
includes a raw data that has not been transformed. In certain
embodiments, the MRI apparatus is for diffusion-weighted MR imaging
(DWI). In certain embodiments, the MRI apparatus is configured for
diffusion tensor MR imaging (DTI). In certain embodiments, the acquiring
of 3-dimensional k-space data includes acquiring substantially entire
3-dimensional data.

[0018] In certain embodiments, the control component includes one or more
devices. In certain embodiments, a single device is configured to
generate the one or more instructions. In certain embodiments, a
plurality of devices are configured to generate the one or more
instructions.

[0019] Certain embodiments of the present disclosure relate to a control
system for an MRI apparatus. The system includes a control component
configured to generate one or more instructions for providing an
excitation radio-frequency (RF) signal, and providing first and second
inversion RF signals to define a field-of-view (FOV).

[0020] In certain embodiments, the first and second inversion RF signals
include first and second inversion RF pulses. In certain embodiments, the
first inversion RF pulse is applied substantially immediately after the
excitation RF signal. In certain embodiments, the first and second
inversion RF pulses are separated by a time approximately 5 ms or larger.
In certain embodiments, the one or more instructions further includes an
instruction for providing slice-selective gradients that are selected
such that magnetization within the FOV is substantially preserved while
magnetization external to the FOV is substantially suppressed, thereby
allowing magnetization in each of a plurality of slices to be
substantially maintained in its equilibrium state while exciting and
imaging one or more others of the plurality of slices.

[0021] In certain embodiments, the control component includes one or more
devices. In certain embodiments, a single device is configured to
generate the one or more instructions. In certain embodiments, a
plurality of devices are configured to generate the one or more
instructions.

[0022] Certain embodiments of the present disclosure relate to a system
for correcting a motion artifact during MR imaging. The system includes a
control component configured to generate one or more instructions for
acquiring navigation data substantially together with imaging data,
determining whether to re-acquire the imaging data based on the
navigation data, and re-acquiring the imaging data based on the
determination.

[0023] In certain embodiments, the motion artifact is due to intra-shot
motion. In certain embodiments, the motion artifact is due to inter-shot
motion. In certain embodiments, the determining and re-acquiring are
performed substantially real-time. In certain embodiments, the navigation
data includes 2D k-space navigation echoes, and the determining includes
identifying value and position of the largest signal in the 2D k-space to
see if either of the value or position is outside of a corresponding
selected range. In certain embodiments, the MR imaging includes a
multi-average singleshot EPI operated as at least one of DWI, DTI, and
fMRI. In certain embodiments, the MR imaging includes at least one of
spin-echo, multiple spin-echo, gradient-echo, and segmented
gradient-echo.

[0024] In certain embodiments, the control component includes one or more
devices. In certain embodiments, a single device is configured to
generate the one or more instructions. In certain embodiments, a
plurality of devices are configured to generate the one or more
instructions.

[0025] For purposes of summary, certain aspects, advantages, and novel
features have been described herein. It is to be understood that not
necessarily all such advantages may be achieved in accordance with any
particular embodiment. Thus, this disclosure may be embodied or carried
out in a manner that achieves or optimizes one advantage or group of
advantages as taught herein without necessarily achieving other
advantages as may be taught or suggested herein.

BRIEF DESCRIPTION OF THE DRAWINGS

[0026] A general architecture that implements the various features of the
disclosure will now be described with reference to the drawings. The
drawings and the associated descriptions are provided to illustrate
embodiments of the disclosure and not to limit the scope of the
disclosure. Throughout the drawings, reference numbers are re-used to
indicate correspondence between referenced elements.

[0047] An improved three-dimensional (3D) singleshot stimulated echo
planar imaging (3D ss-DWSTEPI or ss-STEPI or STEPI) is presented which
includes a novel technique to perform 3D singleshot DWI and DTI of a
restricted 3D volume. In certain embodiments, 3D ss-DWSTEPI acquires 3D
raw data from a limited 3D volume after a single diffusion-prepared
driven-equilibrium (DPDE) preparation by short EPI readouts of several
stimulated echoes. In certain embodiments, the raw data includes k-space
data. In certain embodiments, the raw data includes data that has not
undergone any transformation. The EPI readout time is preferably
shortened by using an inner volume imaging (IVI) technique along the
phase-encoding direction.

[0048] In certain embodiments, 3D ss-STEPI may be used to image any
localized anatomical volume within a body without aliasing artifact and
with high-resolution. The results from 3D ss-STEPI imaging studies of
phantoms, an excised animal heart, and in vivo results from human
volunteers, all demonstrated excellent resolution among all directions.
In certain embodiments, the advantages of STEPI versus existing
techniques for DTI include: (1) no motion-induced artifact, (2) much
reduced susceptibility artifact, (3) high spatial resolution in all
imaging directions, and (4) reduced scan time. STEPI can be also used in
multishot EPI imaging to reduce the total imaging time by a factor equal
to the number of slice-encodings.

[0049] In certain embodiments, a method for interleaved multiple inner
volume imaging (IMIVI) uses two inversion (180 degree) RF pulses to
define the desired FOV. In certain embodiments, the inversion pulses are
applied immediately after the excitation RF pulse and are separated by as
little as about 5 ms. In some embodiments, the inversion pulses are
separated by more than about 5 ms, and in some embodiments, the inversion
pulses are separated by less than about 5 ms. In certain embodiments, the
slice selective gradients applied with the RF pulses are chosen such that
magnetization from within the total volume imaged is preserved, while the
magnetization external to the imaged volume is completely suppressed.
Thus, the magnetization in each slices or slab is only slightly disturbed
from its equilibrium state while exciting an imaging the other slices or
slabs. As a result, the signal loss for all slices is minimal allowing
time efficient interleaved multislice/slab acquisition.

[0050] In some embodiments, real-time navigation (RTN) determines if the
data acquired by any specific EPI readout is corrupted by the subject's
motion, and instructs the MRI scanner to reject and reacquire the data.
2D navigator echoes can be acquired as a part of the data acquisition
process, and then sent to the data processing computer where they are
analyzed by identifying the value and the position for the largest signal
n k-space data of the navigator. If the magnitude or the position of the
peak is out of bounds, the data are rejected and the data acquisition
program is instructed to reacquire the data in real time. RTN can be
useful for any MR imaging techniques, especially for multi-average
singleshot EPI (DWI, DTI, and fMRI). It can be useful for most
conventional MR imaging techniques, which include (but are not limited
to) spin-echo, multiple-spin-echo (fast-/turbo-spin-echo), gradient-echo
(SPGR/GRASS, FLASH/FISP, field-echo), segmented gradient-echo (MP-RAGE).

[0052] As depicted in FIG. 1, DPDE preparation section 110 precedes the
stimulated echo imaging sequence. Two 180° RF pulses are applied
following a 90° excitation pulse. The 180° pulses are used
to determine the localized volume for interleaved multiple inner volume
imaging. The first inversion of the double inversion is used to invert
substantially all magnetization, and ultimately substantially eliminate
unwanted signal from the out-of-volume magnetization. The second
inversion is used to restore the magnetization in other slabs to be
imaged and allows time-efficient interleaved acquisition of multiple
slabs. The earliest group of ACQs between the double inversion and the
first diffusion gradient collects three reference echoes (2 odd and 1
even echo) for EPI phase correction. Before being tipped up to the
longitudinal direction, the diffusion-prepared transverse magnetization
in each voxel is dephased more than 2π by a dephasing gradient
(indicated by right arrow: → in FIG. 1a) to remove the image
intensity dependence upon the tipup RF pulse phase. The residual
transverse magnetization is suppressed by a spoiler gradient applied
after the tipup pulse. The slice-selection gradient (indicated by the
vertical arrow in section (a)) is applied in the slice-encoding direction
for substantially all RF pulses in DPDE preparation, except for the two
IVI refocusing/inversion RF pulses, where the first pulse is spatially
non-selective and the gradient for the second 180° pulse is
applied along the phase-encoding direction to define the reduced phase
FOV.

[0053] IVI limits the excited FOV in the phase-encoding (PE) direction to
include only the anatomy of interest. Time efficient interleaved
multivolume IVI can be obtained because the two refocusing pulses with
slice selection along the phase encoding direction that are used to
create the limited FOV also return most of the out-of-slab magnetization
to the longitudinal direction.

[0054] The data acquisition part of the pulse sequence includes multiple
segments (section (b)). Each segment includes the excitation RF pulse
(creating a single stimulated echo), rephasing crusher gradient, EPI
readout, and rewinding gradients. For each segment, the flip angle of the
imaging RF pulses is gradually increased to reduce the T1
decay-related blurring in the slice direction. The flip angle for the
last segment is 90° to consume substantially all remaining
longitudinal DW magnetization. The rephasing crusher gradient (indicated
by left arrow:  in FIG. 1), applied substantially immediately after
the slice selection gradient of excitation RF pulse, α, rephases
the phase accumulated during the dephasing crusher gradient (→)
prior to the tipup 90° RF pulse. The ETL of the EPI readout in
each segment is chosen to be substantially the same as the number of
acquired ky phase-encodings (e.g., 31 for an imaging matrix with 48
ky views). This number is kept relatively small to reduce
susceptibility artifacts. The phase-encoding order is increased linearly
in each segment, and a center-out slice encoding order is used to improve
the SNR by placing the center of slice-encoding at the earliest echotrain
acquisition. In one example, 62.5% of both phase-encodings and
slice-encodings (ky and kz, respectively) are asymmetrically
acquired to reduce the ETL and the length of the total data sampling
time, i.e., 10 slice-encodings for 16 slices, and the data were
zerofilled and reconstructed using the reconstruction program supplied by
the manufacturer. On completing data acquisition of each echotrain, all
imaging gradients can be completely rewound to preserve the transverse
coherence and maintain some level of steady-state transverse
magnetization.

[0055]FIG. 2 depicts a method of implementing diffusion weighing using 3D
ss-DWSTEPI. The symbols δ and Δ represent the duration of
diffusion gradient and the spacing between two main diffusion gradients,
respectively. An extra delay is inserted between the DW spin-echo
position and the 90° tipup RF pulse, which is same as the time
interval between the center of the imaging RF pulse and the stimulated
echo position. Diffusion weighting may be achieved by applying the
Stejskal-Tanner diffusion-weighting gradient on both sides of the third
180° RF pulse (FIG. 1, section (a)) and additional bipolar
gradients during the delay to maximize the diffusion weighting for a
given TE. Neglecting gradient ramping up/down time, the b value for the
diffusion weighting scheme can be given by:

b=(γGD)2(δ12(δ1-Δ/3)+.d-
elta.23). (1)

[0056]FIG. 3 describes the evolution of spins to form the diffusion
encoded stimulated echoes in STEPI. DW magnetization MD.sup.+({right
arrow over (r)}) is refocused at the spin-echo position "SE". Then,
MD.sup.+({right arrow over (r)}) is dephased more than 2π by the
gradient Gcr1 and the 90°.sub.-x RF pulse tips half of the
dephased magnetization to the longitudinal direction, leaving the other
half in the transverse plane. The transverse component is spoiled by the
spoiler gradient Gsp. As a result the DP magnetization with
dephasing is aligned to the longitudinal direction before the imaging RF.
The imaging RF αn tips a fraction of the magnetization into
the transverse plane, the crusher gradient Gcr2 rephases the
dephasing caused by crusher gradient Gcr1, and a stimulated echo is
formed at the position "STE", at the position where ky=0 in the EPI
readout.

[0057] Eq. (2) describes the longitudinal magnetization just before the
nth imaging RF pulse with respect to the previous longitudinal
magnetization value Mzn-1. Here, αn is the flip
angle of the nth imaging RF pulse and τ is the duration of each
data acquisition segment.

[0058] The two terms in Eq. (2) are the freshly recovered and
diffusion-prepared magnetization, respectively. A signal from the first
term, which is not diffusion-weighted, is spoiled after each excitation
by the rephasing crusher gradient (indicated by left arrow:  in
FIG. 1). As a result, the detected MR signal reflects only the
diffusion-weighted magnetization, yielding a straightforward single
exponential dependence on the applied b value. The diffusion-prepared
longitudinal magnetization decreases along the slice-encoding direction
due to the repeated RF pulses and T1 decay.

[0059] Neglecting the steady-state transverse magnetization, the
diffusion-weighted transverse magnetization after the nth imaging RF
pulse αn can be described as:

MDz({right arrow over (r)},TR,TE,b)=Mo({right arrow over
(r)})(1-e-(TR-TACQ.sup.)/T1.sup.({right arrow over
(r)}))e-bD({right arrow over (r)})e-TE/T2.sup.({right
arrow over (r)})e-TD/T1.sup.({right arrow over (r)}), (4)

for diffusion weighting b, effective echo-time TE=TE1+TE2, a
time delay, TD, between the tipup RF pulse and the first imaging RF pulse
(α1), and total pulse sequence duration TACQ which
includes the diffusion-preparation and complete 3D data readout. The
factor 1/2 arises from applying the pre-tipup dephasing gradient to
remove the signal dependency on the relative phase between the tipup RF
pulse and DW magnetization (20).

[0060] As shown by the second term in Eq. (2), the measured signals
experiences T1 rather than T2 decay along the slice-encoding
direction. This can be very advantageous, because T1 is typically an
order of magnitude longer than T2 in most tissues. Blurring in the
slice-encoding direction, which may arise from T1 decay during the
long data acquisition, may be reduced by using variable flip angles. The
transverse magnetizations Mn-1.sup.+({right arrow over (r)},t) and
Mn.sup.+({right arrow over (r)},t) after two consecutive RF pulses
(αn-1 and αn) are:

[0061] To achieve equal signal amplitude (Mn.sup.+({right arrow over
(r)},t)=Mn-1.sup.+({right arrow over (r)},t)) the relationship
between the flip angles of two adjacent RF pulses should satisfy

tan αn-1=sin αne-τ/T1.sup.({right
arrow over (r)}) (6)

[0062] The flip-angle for the last RF pulse is set to 90° to
consume substantially all remaining longitudinal magnetization, and the
flip angles of the proceeding RF pulses can be calculated using the
relation in Eq. (6), and typical values of τ and T1 which are
about 40 ms for 31 ETL with the receiver bandwidth of about 1.086
kHz/pixel and T1=1.0 s (approximately) for white matter at 3T. In
general, the use of this equation resulted in very small initial flip
angles and, correspondingly, low SNR images. For this reason a compromise
was made between T1 decay related blurring in the slice encoding
direction and image SNR, using a ramped variable flipangle with a larger
starting angle and smaller increases, ending again in a 90° pulse
to consume substantially all remaining longitudinal magnetization.

[0063] After completion of each EPI echotrain, the remaining transverse
magnetization can be rewound or completely spoiled by a spoiler gradient.
Because the central planes of k-space are acquired during the first few
stimulated echoes, with rewinding or with spoiling the DWI signal
intensity undergoes simple exponential decay with respect to the b value,
as:

[0064] FIG. 4 depicts images obtained by 3D ss-DWSTEPI for reduced phase
FOV preparation. The prescribed volume is indicated by the dotted box in
section (b). The resulting images have a good quality without any
aliasing along the phase-encoding direction from the phantom regions
external to the prescribed volume. This result demonstrates the
applicability of our reduced FOV preparation scheme to limit FOV in the
phase-encoding direction. Such a restricted in-plane FOV can be sampled
by a short EPI readout resulting in significantly reduced image
distortion due to local magnetic field susceptibility. The degree of
distortion in the images acquired by 3D ss-DWSTEPI was comparable to that
of 2D ss-DWEPI with analogous EPI readout duration. Note that the 3D
ss-DWSTEPI data acquisition was accomplished with 15 applications of the
excitation RF pulse followed by the EPI acquisition of 31 gradient
echoes. The duration of each segment including the RF pulse and complete
ky acquisition of 31 gradient echoes was about 38 ms. The total
duration of the 3D data acquisition was around 560 ms for the 15 actual
slice-encodings required to reconstruct a 24-slice volume.

[0065] In certain embodiments, 2D STEPI has been tested and given MR
images with greatly reduced geometric distortion. It may be an important
tool for a high-field MRI system, such as a system where b of equation 7
is greater than or equal to 3T. In 3D STEPI, kz (slice) encoding may be
segmented along a stimulated echo train, and each segment can complete
the entire ky (phase)-encoding. In 2D, the total ky is
interleavely segmented into multiple stimulated-echotrains. The geometric
distortion of the resultant MR images is 1/(Number of segments) of that
of conventional 2D singleshot-EPI.

[0066] As shown in FIG. 5, the signal loss in interleaved multislab
imaging with reduced FOV preparation was substantially reduced using the
new technique with double inversion (quadrature) compared to the rapid
signal decay for the standard method with a single inversion pulse
(Δ). Adiabatic RF pulses with about 5.12 ms duration were used to
implemented double inversion reduced FOV preparation. The separation
between the inversion RF pulses in the new technique was around 6.0 ms.

[0067]FIG. 6 illustrates two sets of 12 slices from 4 contiguous 12-slice
slabs. Two separate acquisitions (passes) were used to image slabs 1 and
3 in the first acquisition, and slabs 2 and 4 for the second. The signal
loss evident on the edge slices of each slab is due to the RF profile
variation and is a common problem in most 3D imaging methods.

[0068]FIG. 7 depicts the affects of decreasing the number of
slice-encodings. More particularly, by reducing the number of
slice-encodings, the blurring in the slice direction can be reduced.
Blurring can also be improved by using a small starting flipangle with an
increased number of averages to obtain acceptable SNR. The peak
amplitudes of the stimulated echoes are plotted section (a), with respect
to their occurrences in time relative to the excitation RF (t=0) in DPDE.
The amplitude of the later echoes of 3D data readout was about 40%,
compared to the first echo (kz=0). The corresponding
point-spread-function (PSF) is shown in section (b). The full width at
half of maximum for the PSF was about 1.8 pixels indicating that image
blurring in the slice direction was mild. The 3D interleaved multiple
inner volume ss-DWSTEPI images shown in FIG. 4 (xy-plane) and section (c)
of FIG. 7 (xz-plane) demonstrate high resolution without any noticeable
blurring in the slice encoding direction.

[0069] The image distortion observed in STEPI is a function of the number
of echoes in the EPI echotrain. Because there are typically more phase
encodings than slice encodings, the number of echoes in the EPI echotrain
can be reduced by interchanging phase and slice encoding. With this
switch, slice encoding is performed in conjunction with the EPI readout
and one phase encoding is applied for each EPI segment.

[0070] The results of a DTI study of a canine heart ex vivo are shown in
FIG. 8. There are some residual aliasing artifacts at the left portions
of the images along the phase-encoding direction. The helical structure
of the myocardial muscle is well presented in the color map, similar to
results previously reported from the excised animal hearts.

[0071] The images from a DTI study of the midbrain of a healthy volunteer
are presented in FIGS. 9-11. Note that the images were acquired using a
head coil, in which the signal reception sensitivity rapidly drops near
the mid level of cervical spinal cord. FIG. 9 shows 9 central slices from
16 contiguous slices covering an approximately 20 mm thick slab. The
bright signal indicated by the arrow appears to be a susceptibility
induced artifact.

[0072] DW images of the central slice are shown in FIG. 10 for b=0
s/mm2 and about 400 s/mm2 for 7 non-collinear directions. DW
images were processed to estimate DTI parameters, such as eigenvectors,
eigenvalues, and fractional anisotropy (FA). The resultant FA maps and
RGB colored maps of the principal eigenvector are presented in FIG. 11
for the central 9 slices, which completely cover the cervical spinal cord
in the transverse direction. These results are very promising for in-vivo
human applications of 3D ss-DWSTEPI for high-resolution DTI.

[0073] 3D ss-DWSTEPI can acquire the diffusion-weighted magnetization of a
localized volume after a single diffusion preparation. Even though
spatial coverage is limited in the phase and slice directions, the FOV in
the readout direction can be arbitrary and limited by the desired image
dimensions and the sensitivity volume of the receiver coils. The system
and method described herein not only reduces susceptibility artifacts by
using significantly shortened EPI readouts, but also freezes most of the
physiologic motion by using a single short data acquisition. 3D
ss-DWSTEPI can be useful for high resolution 3D DTI of limited volumes of
interest such a localized brain regions, cervical spinal cord, optic
nerve, heart or other extracranial organs.

[0074] Subject motion during the diffusion gradients can cause shading and
ghosting artifacts in the resultant diffusion-weighted (DW) images [1-3],
and consequently result in reduction of the accuracy of DTI measurement
in multishot DW imaging. Typically, DW images are acquired with multiple
signal averages to improve SNR. Any discrepancy in subject position
between averages would result in blurring of the averaged image. We
hypothesized that any motion might decrease the accuracy of DTI
measurement in multishot or multiple averaging singleshot DTI. In this
report, a motion artifact correction scheme for multi-shot 3D EPI-DTI
with one EPI readout per kx-ky or kx-kz plane is described. This
technique can be used to reduce artifacts caused by subject motion during
diffusion gradients or subject motion between shots or averages and,
therefore, to improve the accuracy of DTI measurement.

[0075] Data acquired by multi-shot DTI pulse sequences can be corrupted by
two principal types of motion: motion during diffusion gradient
application (intra-shot motion) and motion between shots (inter-shot
motion). Intra-shot motion can cause a significant signal loss when even
small motion can cause partial or complete dephasing or an additional
phase factor in image space. Inter-shot motion caused by global changes
in subject position results in additional phase term in k-space data
(translation) or k-space data shift (rotation). Both types of motion can
be identified and corrected if multi-dimensional navigators are acquired
together with the imaging data. A 3D multi-shot EPI-DTI pulse sequence
with a limited FOV preparation and 2D navigators was implemented. In the
sequence, two techniques were used to resolve the motion: (1) real time
(RT) navigation of data acquisition and (2) correction of inconsistencies
between shots using 2D navigators. The first technique, RT navigation,
can be used to monitor/identify the shots corrupted by substantial
motion, and direct the pulse sequence to reacquire data for those shots.
The second method can identify and remove inconsistencies caused by small
motions between shots and can identify and correct the resulting subject
position changes. The method for motion artifact correction in 3D
multi-shot DTI with 2D navigator echoes is schematically described in
FIG. 12.

[0077] The input data for the technique are the dataset with excessive
motion corrupted echoes reacquired in real-time. The measurement data
were obtained using 3D multi-shot EPI-DTI pulse sequence with limited FOV
preparation. The imaging parameters were: b=500 sec/mm2,
TR/TE=4000/75 ms, ETL 33, 192×33×8 imaging matrix, and 4
averages. The first average was treated as the acquired data, while the
remaining 3 averages were considered as the reacquired data to simulate
RT navigation. The agar phantom was intermittently moved predominantly in
the vertical or horizontal direction during the acquisition to mimic
physiologic motion. The algorithm described in FIG. 12 was used to
correct the phase inconsistency between shots. First, 2D navigator echoes
Nkz(kx,ky) were Fourier transformed and the corresponding
2D phase maps φkz(x, y) were constructed. These maps were
combined with the associated imaging echoes as I'(x, y, kz)=I(x, y,
kz)eiφkz.sup.(x,y) [1]. Inter-shot in-plane motion
(translations (Rkzx, Rkzy) and rotation
(θkz) in x-y plane) can be estimated from the 2D navigator
image, Akz(x,y) using the methods described in [3]. The corrected
dataset was used for image reconstruction.

[0078] FIG. 13 illustrates the real-time reacquisition of the imaging
echoes corrupted by excessive motion. Numbers on top of the figure
represent shot number and the numbers with prime (4' and 3') represent
the reacquired shot. The phantom was (a) stationary or intermittently
moved/rotated during the acquisition mainly in (b) vertical or (c)
horizontal direction. The RT navigation directed the pulse sequence to
reacquire the motion-corrupted echoes. The shot where the peak of the
associated 2D navigator k-space data is shifted noticeably were replaced
by the corresponding shot from next average (RT navigation). The
resulting dataset was used for the phase-correction.

[0079]FIG. 14 shows the images reconstructed from the original dataset
(row 1), the dataset after RT navigation (row 2), and the dataset after
phase inconsistency correction (row 3). The images from row 1 in column b
and c are corrupted by motion artifact. The reacquisition of data by RT
navigation improves image quality significantly (row 2 of FIG. 14) but
leaves some residual ghosting in the slice direction, especially in
images corrupted by vertical motion. These residual artifacts are removed
after phase correction (row 3 of FIG. 14). The horizontal banding in all
images is a systematic artifact arising from the shape of the RF profile
used for reduced FOV preparation and does not affect the DTI processing
as it presents in all images.

[0080] With reference to FIG. 14, the phantom was stationary or
intermittently moved vertically or horizontally during data acquisition.
The DW images in the second row were reconstructed from the dataset where
the motion corrupted EPI readouts were replaced by the corresponding
readouts from next average dataset (simulation of RT navigation). The
dataset after RT navigation was additionally corrected using 2D navigator
phase information. The phase correction combining with the RT navigation
gave substantially improved DW images.

[0081] In some embodiments, a new imaging technique, 2D singleshot
Real-Time Navigated diffusion-weighted EPI (2D ss-RTN-DWEPI), is used to
acquire DWI, which the data is monitored in real-time by using the
real-time feedback capability of the MRI system, to identify the data
with substantial corruption due to the motion. The largest echo peak of
2D navigator echoes is searched and its magnitude and the position
calculated for evaluation of RTN test by comparing the later averages to
the first average of the same diffusion encoding in real-time. If the
differences are out of bound from the given threshold, the data is
rejected and immediately reacquired in real-time.

[0082] The RTN data acquisition is implemented into 2D ss-EPI, using the
IDEA pulse sequence development environment, and the real-time
calculation of the navigator data is implemented into an image
construction program in Image Construction Environment (ICE) (Siemens
Medical Solutions, Erlangen, Germany). The pulse sequence (FIG. 15) is
capable of Interleaved Multi-slice Inner Volume imaging (IMIV) of a
reduced field-of-view (FOV) without the aliasing artifact in
phase-encoding direction, using the application of the double inversion
immediately after the 90° excitation RF. An adiabatic RF pulse is
used for double inversion for IMIV. The first inversion RF does not
accompany the slice-selection gradient and the second with
slice-selection gradient in phase-encoding direction that defines the
reduced FOV phase-encoding direction. The application of double inversion
increases TE by about 12 ms, which includes two 5.12 ms RF pulses and two
pairs of crusher gradients, sandwiching around the RF pulse to destroy
the free-induction-decay caused by imperfect 180° pulse which
generally induces the stimulated echo artifact.

[0083] Diffusion weighting is accomplished by adding a pair of
Stejskal-Tanner diffusion-weighting gradient on both sides of the
refocusing 180° RF pulse (the third 180° RF in the
diagram). Real-time feedback flag (RT_FEEDBACK) is added to all EP
imaging echoes. Therefore, the acquired data serves as the 2D navigator
echoes as well as the imaging echoes. Upon the transfer of the raw data
to the image construction computer, the echo data with the flag
RT_FEEDBACK is immediately fed into the real-time calculation algorithm,
which searches for the largest echo in k-space and calculated its
coordinate kpk and magnitude m(kpk) and sends these values to
the scanning computer for the RTN test. The location kpk of the
largest peak of the 2D echoes corresponds to the center of k-space.
Images are constructed for the shots, on which the RTN test is passed.

[0084]FIG. 15 illustrates a pulse sequence diagram of 2D ss-RTN-DWEPI
with interleaved multislice inner volume imaging. Vertical arrows
indicate the slice-selection gradients. Two 180° RF pulse enclosed
within the dotted box are for reduced FOV imaging in phase-encoding
direction. Slice selection gradient is applied in phase-encoding
direction to restrict the imaging FOV. EP echoes were also set for
RT_FEEDBACK flag for the real-time calculation. The dashed vertical line
indicates the position of spin-echo and the center of k, i.e., k=0.

[0085] One may assume that a human subject is generally motivated to hold
still at the earlier acquisition of imaging than the later ones. The
first averaging data for all diffusion encoding directions are acquired
at the early acquisition by using the long-term average mode, which the
averaging resides at the outermost loop of the data acquisition. The
first average data of each diffusion encoding direction was then
considered as the reference shot without motion corruption. The scanning
computer stores the reference values (m(kpk; kpk) of the first
averages to compare those of the later shots to monitor the change of the
magnitude and the position of the largest echo in k-space and to
determine if the data are acceptable.

[0086] The flowchart shown in FIG. 16 describes the RTN procedure of 2D
ss-RTN-DWEPI. The dashed and the dotted boxes indicate the processes in
the image acquisition and the image construction computers. The real-time
calculation is processed by ICE program within the image construction
computer. Indices na, nreacq, and dj indicate average,
reacquisition, and diffusion encoding counts, respectively. The position
kpk,1, and the magnitude m1(kpk,1), of the largest echo in
2D navigator echoes of the first averaging shot (na=1) are stored
into a temporary memory as the reference values and are used for the
comparison for other averaging data. For a specific diffusion encoding
direction dj and averaging count na, the maximum reacquisition
is set to avoid the extended acquisition duration. If the data
successfully pass the RTN test or the reacquisition count reaches the
maximum, the reacquisition count nreacq is reset to zero, which also
directs the image construction computer to construct the images and send
them to the image database. The averaging proceeds to the next
acquisition.

[0087] In FIG. 16 thick arrows represent the data transfer between the
acquisition and the reconstruction computers. All NMR data which is also
2D navigator echoes are sent to the reconstruction computer to search for
the largest echo peak and to calculate its magnitude and the position in
k-space. These values are transferred to the scan computer to determine
if the data is acceptable compared to the first repetition data.

[0089] Imaging Phantom: A cylindrical phantom filled with the mixture of
water and agarose was used. T1 of the phantom was measured about 2.0
s at 3 T. The phantom was intermittently lifted from one end or
horizontally rotated during the acquisition. The imaging was accomplished
using the parameters, the receiver bandwidth of 1.086 kHz/pixel and 31
actual echoes per EPI echotrain. A single channel transmit/receive RF
coil was used for the simplicity of the raw data. Other parameters
typically were that slice thickness was 2.0 mm, and TR/TE was 4.0 s/60
ms. The imaging matrix was 128×48 and 16 slices, with 62.5%
asymmetric acquisitions in phase-encoding direction, which covered 96 mm
and 32 mm in the phase and slice encoding directions for about 2.0 mm
isotropic resolution. For the phantom imaging, 37.5% of full FOV was
imaged in phase-encoding direction, using IMIV to image the reduced FOV
without aliasing artifact. The averaging was accomplished via the
magnitude averaging to remove the motion artifact. Otherwise the phase
instability among the different averages may deteriorate the image
quality.

[0090] A set of 2D ss-RTN-DWEPI images was acquired with diffusion
weighting b=0 and 500 s/mm2 in four non-collinear directions,
(1,0,0), (0,0,1), (1,0,1), (-1,0,1) in physical gradient coordinate
(Gy, Gx, Gz) that represent (vertical, horizontal, magnet
bore) or anterior/posterior, right/left, superior/inferior (A/P, R/L,
S/I) anatomic directions. The phantom was intermittently moved during the
acquisition to simulate a subject's random physiological motion. The
maximum number of reacquisition was set to 2 and the threshold values for
RTN testing were 30% and 2Δkx for changes of the magnitude and
the peak position. The minimum number of slices for the failure of RTN
test was set to 2 for the interleaved multislice 2D imaging. For
instance, if two or more slices out of total slices would fail RTN test,
the acquisition of whole slices were repeated for that average that
increased the total imaging time by an additional TR.

[0091] Human Imaging: To demonstrate the feasibility of the RTN for human
imaging, 2D ss-RTN-DWEPI was applied to acquire DWI from human volunteers
using image matrix 128×96 with an FOV of 256×192 mm, TR/TE
5.0 s/66 ms, b of 0 and 750 s/mm2 in 7 non-collinear directions
along three orthogonal axes and four tetrahedral vertices: (1,0,0),
(0,1,0), (0,0,1), (1,1,1), (-1,-1,1), (1,-1,-1), (-1,1,-1). A twelve
channel receive-only head-matrix coil (Siemens, Erlangen, Germany) was
used. The channel within the posterior matrix was selected for RTN
evaluation. During the data acquisition, subjects were instructed to move
the head along anterior-posterior direction. Due to the narrow space in
the headcoil, the subject's motion was mostly the rotational motion. DW
imaging was repeated for the same volunteer without head motion. The
subject was instructed to hold the head position, but with free breathing
and swallowing. The imaging protocol was approved by the University of
Utah Institutional Review Board and the informed consents were collected
from volunteers.

[0092] FIGS. 17a-17c illustrate motion corrupted magnitude and the
corresponding phase images of the slice 7 in different diffusion encoding
directions are compared with respect to motion free shot. Phase-encoding
was in horizontal direction. The solid and the dotted circles represent
the accepted and the rejected measurements, respectively. The numeric
numbers enclosed within the rectangular boxes represent the different
slices in a given repetition time TR. The numbers enclosed within the
circle in FIG. 17a and FIG. 17c indicate the different diffusion encoding
directions, which includes 0 for b=0 s/mm2, and those in the
rectangular boxes are for the different slices. The data, which failed
the RTN test and therefore is rejected, is indicated by the number
encircled by the dotted lines, and the accepted data are enclosed by the
solid circles. From Fourier transform theory, one pixel shift in k space
corresponds to 2π phase difference on the edge of FOV in image space.
The number of the wraps in phase images represents the number of shift of
Δkx in k-space. The shots that either the magnitude of the
navigator changed more than 30% or the k=0 peak shifted more than
2Δkx(=2/FOVx) from the reference values were reacquired.

[0093] The measured raw data was transferred from the scanner to a
computer for further analysis. The magnitude image for diffusion encoding
direction 2 indicated substantial reduction compared to other directions
(3 and 4). This shot was rejected and reacquired in real-time. The
reacquired image encircled by solid line for direction 2 demonstrated no
motion artifact.

[0094] For direction 3, although the first attempt for slice 7 succeeded
for the RTN test as in the figure, RTN reported the failure on this
measurement, because other slices (S1and S3) failed the test as
indicated in FIGS. 3b and 3d. k=0 peaks of the slices 1 and 3 shifted by
3Δky and 2Δky respectively. Therefore, the whole
acquisition of direction 3 was repeated.

[0095] FIGS. 18a and 18b illustrate plots of the echo-peak position in
k-space along various diffusion encoding directions of b=0 and 400
s/mm2. Numbers in horizontal axis represent b=0, (1,0,0), (0,1,0),
(0,0,1), and (1,1,1). The shift was plotted separately along the readout
and phase-encoding directions. The change of the magnitude and the peak
shift of the largest echo in 2D navigator are plotted in FIGS. 18a and
18b, respectively. The numbers in horizontal axis represent the diffusion
encoding directions, which includes 0 for b=0. The shift of the echo peak
was calculated from the peak position of the reference navigator echo.
When the phantom was moved along phase-encoding direction, the peak shift
occurred in ky direction, while it was observed in kx or the
motion along the readout direction. As indicated in the direction 2 in
FIGS. 17a and 17c, the magnitude was decreased by 22.5% and k=0 peak
shifted about 11 units of Δk in phase-encoding direction.

[0096]FIG. 19 illustrates the magnitude and the phase images of (a, d)
the reference (the first average), (b, e) the motion-corrupted, and (c,
f) the reacquired data in DW imaging of a human volunteer. The volunteer
intentionally nodded his head to initiate the rotational motion during
the acquisition of corrupted shot. The corrupted and reacquired images
clearly indicated the significant drop of the magnitude in FIG. 19b and
the improvement in FIG. 19c, respectively. RTN test reported 52% change
of the magnitude and 7 Δk shift of the navigator peak for this
specific data in FIG. 19(b, e) compared to the first averaging data in
FIG. 19(c, f). This shot failed both magnitude and shift tests.

[0097] The motion artifact in FIGS. 19a, b is an exceptional case of
motion, which was caused by intentional head motion, and which will be of
a greater magnitude than that normally encountered when imaging a patient
endeavoring to remain still. Because the source of motion artifact for
most of corporative human subjects may be breathing and swallowing
motions, all others data were much cleaner than these images. The images
in FIG. 20 are from a volunteer, who stayed still during the entire DW
imaging acquisition except a few swallowing motion. The images in FIG. 20
are the magnitude and the phase images of (a, d) the reference (the first
average), (b, e) the motion-corrupted, and (c, f) the reacquired data.
RTN process reported the magnitude change below 0.5% and 2 Δk shift
of the navigator peak echo. This shot succeeded the magnitude test,
however it failed for the phase test and the entire slices were
reacquired in real-time.

[0098] Intra-shot motion particularly during the application of the
diffusion gradients can cause a significant signal loss when even small
motion can cause partial or complete dephasing or an additional phase
factor in image space. The motion caused by global changes in subject
position results in additional phase term in k-space data (translation)
or k-space data shift (rotation). Both types of motion can be identified
by using the navigator data.

[0099] Once the motion corruption occurs within the DW images, it is
difficult to correct the artifact. As described herein, 30% for magnitude
variation and 2 Δkx for the peak shift of k=0 point in k-space
were used for the acceptable criterions. Either/both decreasing the
magnitude change or/and increasing the peak shift will increase the
imaging time, which can induce the increased chance of the position
change among the shots. Limits on the number of reacquisition of a
specific set of measurements can be specified to 2 for imaging of both
the phantom and the brain of human volunteer to avoid unacceptably long
acquisitions. It may be increased to image the subjects who may be
somewhat corporative. However the increased maximum reacquisition may
extend the total imaging time for incorporative patients, because there
may be frequent failure of RTN test for all diffusion encodings. The
bounds for these values may be based on the ranges that yield minimal
artifact and minimal reduction in DTI accuracy in minimal increase of the
acquisition time.

[0100] During the preliminary imagings, there were occasions which the
maximum reacquisition count was consumed without passing the RTN test and
the acquisition proceeded to the next acquisition. The acquisition with
the least change of the magnitude and the position of the largest echo
peak was selected for averaging in offline, using the measurement raw
data. This algorithm may be implemented into the online reconstruction
program.

[0101] If the subject moves and the data are corrupted while acquiring the
first average, all subsequent repetitions may fail to pass the acceptance
criteria. Then the acquisition may be stopped and the subject may be
instructed to hold still and the acquisition may be restarted.

[0102] Upon using multi-channel receive-only coil, the channel closed to
the region of interest was selected for RTN evaluation. For instance, a
channel in the posterior matrix element can be selected for DTI of
cervical spinal cord. For multi-slice imaging, the comparison is made
slice-by-slice. If the navigator echoes for given number of slices cannot
satisfy the acceptance criteria, this shot is reacquired.

[0103] There is a minimum duration for RTN process which includes the
real-time data communication between the acquisition and the
reconstruction computers and the calculation in the reconstruction
computer. This delay must be increased with the increased complexity of
the calculation, such as including Fourier-transformation. It reduces the
maximum number of slices for a given TR. 1 ms was long enough for current
study because the process in real-time ICE program was simple.

[0104] If the motion of the imaging subject caused the change of position
from previous averages and occurred between the shots, RTN would not be
able to detect because the there would not be significant change in the
magnitude and in the position of the navigator echo peak. The RTN
algorithm may be modified such that the Fourier-transformed magnitude
image is subtracted from the reference image and the total signal of the
difference image is summed up and used to detect the change of the
position which may have occurred in between the current and previous
shots. The RTN may not reject these data; rather a postprocessing may be
used to co-register the later averages with respect to the early ones
before the magnitude averaging.

[0105] In multishot diffusion MRI, navigator echoes are used to correct
the instability of the phase error among the data for different segments,
which include the self-navigating techniques such as SNAIL
(self-navigated interleaved spiral) and PROPELLER (periodically rotated
overlapping parallel lines with enhanced reconstruction) that use the
imaging echoes as the navigator to directly monitor the acquired data.
These imaging techniques use the navigator echoes in post-processing, not
in real-time. Real-time navigation has been used to directly measure the
fat signal within the FOV in cardiac imaging. The RTN technique can be
implemented into a multishot DWEPI sequence by acquiring an addition
echotrain that samples 8-16 echoes of the center of k-space.

[0106] RTN imaging is more suitable to identify global rather than local
motion. It can be also used to detect and monitor voluntary local motion
such as swallowing, which induced the shift of the navigator peak by a
few Δk, as demonstrated in images in FIG. 20. It may be acceptable
for DTI of brain. However, swallowing may induce the anterior-posterior
motion on DTI of the cervical spinal cord. Because the swallowing does
not happen frequently, the RTN parameters may be set to screen any motion
with equal or larger amplitude than the swallowing. If the selected coil
segment is sensitive to a local motion, such as a CSF pulsation for DWI
of cervical spinal cord, RTN test may fail more often that is desirable.
In these cases however, the RTN technique may be combined with cardiac
gated acquisition for diffusion-weighted imaging. Since each
reacquisition increases the total imaging time by TR, the threshold
values for magnitude change and the peak shift of the 2D navigator echoes
for RTN may be increased to detect the data corruption due to the large
motion only and to reduce the number of failures in RTN test.

[0107] As a result, real-time navigated data acquisition for
diffusion-weighted imaging improve the DTI measurement result by
identifying and reacquiring the data with excessive motion-related
corruption in real-time. It can be used to reduce the inconsistency among
the different averaging caused by subject motion during the application
of the diffusion gradients and/or between shots and, therefore, improves
the accuracy of DTI measurements. This technique can be particularly
useful to detect the global motion of the imaging region.

[0108] While certain aspects and embodiments have been described, these
have been presented by way of example only, and are not intended to limit
the scope of this disclosure. Indeed, the novel methods and systems
described herein may be embodied in a variety of other forms without
departing from the spirit thereof. The accompanying claims and their
equivalents are intended to cover such forms or modifications as would
fall within the scope and spirit of this disclosure.